This invention relates generally to methods and apparatus for detecting radiation in CT imaging and other radiation imaging systems, and more particularly to scintillator arrays having increased reflectivity.
In at least some computed tomography (CT) imaging system configurations, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the “imaging plane”. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile.
In some known CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal spot. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator adjacent the collimator, and photodetectors adjacent the scintillator.
One or more rows of scintillator cells or scintillator pixels are provided in a detector array configured to acquire projection data from which one or more image slices of an object are reconstructed. One known detector array includes a two-dimensional array of scintillator cells, with each scintillator cell having an associated photodetector. An epoxy material is used to cast the scintillator cells into a block having specified dimensions for easier handling. To maximize reflectivity and to prevent cross-talk between adjacent detector cells, the cast reflector mixture includes a material having a high refractive index, such as TiO2. Thus, light generated in the scintillating material by impinging x-rays is confined to the detector cell in which it is generated. However, neither the epoxy, the TiO2, nor their mixture are particularly absorptive of x-rays. Thus, neither the photodetectors nor the cast reflector mixture itself is protected from damage caused by impinging x-rays.
In one known cast reflector mixture, a small amount of an oxide of chromium is also incorporated in the cast reflector mixture to further reduce cross-talk between cells. However, inclusion of this material reduces the efficiency of the detector, because the absorbed portion of the generated visible light is never detected by the photodetectors. Chrome is used to reduce the cross-talk by absorbing the light transmitting through the cast walls between pixels. This dopant significantly reduces the reflectivity from 98% to 82%, thus leading to very low light output. The introduction of chrome can reduce the light output by as much as 60% or higher. Also, the ability to reduce the cross-talk is limited. Newer CT applications require higher and higher resolution which means that the new design will require smaller pixels. The light output is further reduced as the pixel size become smaller because the relative opportunity of losing light on the surface is higher. The light output will be reduced by 20 to 25% when the pixel size is reduced by 50%. The low light output can cause image quality problems because of the low signal to noise ratio.